Oxygen-generating biodegradable surgical mesh

ABSTRACT

The present disclosure relates to a novel oxygen-generating biodegradable surgical mesh, and to methods of making and using the novel oxygen-generating biodegradable surgical mesh. More specifically, a novel surgical mesh has been developed, wherein the surgical mesh has a flexible basic structure and comprises a plurality of pores, wherein the surgical mesh has a first face and a second opposite face, wherein the surgical mesh is made of a substantially homogeneous material comprising a biodegradable polymeric material and an oxygen-generating material.

TECHNICAL FIELD

The present disclosure relates to a novel oxygen-generatingbiodegradable surgical mesh, and to methods of making and using thenovel oxygen-generating biodegradable surgical mesh.

BACKGROUND

This section introduces aspects that may help facilitate a betterunderstanding of the disclosure. Accordingly, these statements are to beread in this light and are not to be understood as admissions about whatis or is not prior art.

Over 20 million hernia repairs are performed every year, costing theworldwide healthcare system over $94 billion. Surgical mesh is commonlyused as a structural support for the affected tissue in hernia repairsurgeries in addition to supplying a scaffold for tissue to grow intoand form scar tissue, closing the hernia defect. Mesh is used inapproximately 90% of inguinal hernia repairs, which account for about75% of all hernia cases. While the use of mesh in hernia repair willdecrease recurrence rate, there are some complications associated withthe use of surgical mesh, including increased risk of infection,adhesion, and bowel obstruction. These complications may be due to theproperties of the mesh being used, surgical conditions, or potentiallyprolonged hypoxic conditions slowing the rate of healing.

Researchers have developed many different types of meshes in order tofind the “ideal” surgical mesh, which provides adequate long-termtensile strength, flexibility, biocompatibility, and proper integrationinto the tissue. These meshes can be broadly placed into two categories:resorbable and non-resorbable. Non-resorbable polymeric meshes, mostcommonly made from polypropylene, have a very low recurrence rate, butremain in the body permanently and are more likely to cause chronicpain, limited flexibility at the surgical site, and can erode into thesurrounding tissue.

Resorbable meshes that are currently available do not maintain theminimum tensile strength of 24 kPa for the time needed for the surgicalsite to heal completely. Due to the lack of sustained structuralsupport, a higher recurrence rate is reported for these resorbablemeshes. In addition to higher recurrence rates, resorbable meshes aresometimes made of biological material which can be up to 100 times moreexpensive than synthetic, non-resorbable meshes. These biological meshesare derived from decellularized tissue to form collagen scaffoldscreated from porcine, human or bovine sources. Biological meshes don'ttrigger foreign body responses and provide better integration into thehost tissue, but due to higher recurrence rates and higher costs, thecommercial impact of these meshes is limited. A synthetic resorbablemesh with a longer life span would be advantageous as it would allow forunique processing methods and utilize higher performance materials suchas resorbable polymers or composites to create a product that would notpermanently remain in the body as a foreign object and still providesthe required structural support.

Synthetic mesh has historically been made through one of the fewprocessing methods: knitting, weaving, or expanded polymeric networks.Woven meshes require tightly packed filaments in order to avoid shiftingof filaments and as such are largely not used in clinical applications.Knitting leads to a generally more flexible material, but it isgenerally weaker and also has a high variability of mechanicalproperties depending on the direction of applied stress. Expandedpolymeric networks utilize a network of extremely fine interconnectedfibers to form a sheet but can create pores too small for white bloodcells. Bacteria are typically smaller than white blood cells; therefore,if pores are too small for white blood cells, bacteria would be able toproliferate in the pores and potentially lead to an infection.

Previous studies have investigated the potential to alleviate the risksassociated with infection, adhesion to the surgical site, and preventforeign body reactions to the mesh by coating the meshes withantibacterial materials (i.e. silver nanoparticles) or with a variety ofother materials including collagen and omega-3 fatty acids. While thesecoatings are shown to prevent some of the aforementioned complications,most of the previous works have focused on coating of non-resorbablematerials, hence they have the drawbacks of permanently remaining in thebody, causing chronic pain, and limited flexibility.

When the surgical mesh is implanted into the body during a surgicalprocedure, the vascular structure of tissue is interrupted. The vascularsystem is responsible for bringing oxygen and other nutrients throughoutthe body, therefore when it is damaged, it limits the amount ofnutrients locally available in the body. When normal oxygen levels arenot present in the body, the area can progress to ischemic conditionsthat result in cell death and slowing down the healing of the wound.Therefore, many researchers have conducted studies to develop methods ofgenerating oxygen locally in a wound to aid in the healing process. Acommon strategy to generate oxygen is through an implantable device thatgenerates oxygen through electrolysis or chemical reactions. Whileelectrolytic devices are able to produce oxygen for extended periods oftime, they often require electronic components that remain in the bodypermanently or generate harmful compounds, such as chlorine, asbyproducts. Oxygen generating compounds, however, are able to generateoxygen over the required timeframe while being absorbed into the bodyand avoiding harmful byproducts.

Therefore, there is an unmet need for a novel and economic surgical meshthat may address the issues.

SUMMARY

The present disclosure relates to a novel oxygen-generatingbiodegradable surgical mesh, and to methods of making and using thenovel oxygen-generating biodegradable surgical mesh.

In one embodiment, the present disclosure provides a surgical mesh,wherein the surgical mesh has a flexible basic structure and comprises aplurality of pores, wherein the surgical mesh has a first face and asecond opposite face, wherein the surgical mesh is made of asubstantially homogeneous material comprising a biodegradable polymericmaterial and an oxygen-generating material.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates the potential application of the proposedbiodegradable and oxygen generating surgical mesh, CPO particles areidentified as black dots in the zoomed in portion of mesh.

FIG. 2 illustrates the fabrication process for CPO-PCL mesh. (A) PCLpellets are hot melt mixed with CPO. (B) This mixture is compressionmolded into a 150 μm thick film (C), then laser processed (D) into thefinal thin film mesh. (E) Schematic representation of mesh withhoneycomb pores.

FIG. 3 illustrates (A) From left to right: PCL pellets, CPO powder andcooled mixture of PCL and CPO. (B) Thin film PCL-CPO (CPO 5 wt %, topleft), and two pieces of flexible laser cut meshes. The mesh on theright shows the flexibility of the mesh.

FIG. 4 illustrates thermogravimetric analysis (TGA, 4A) and differentialscanning calorimetry (DSC, 4B) of pure PCL, pure CPO, and the 15 wt %CPO thin film.

FIG. 5 illustrates oxygen generation measurements of water and all thinfilm samples over 17 hours. Thin film samples were tested in phosphatebuffered saline (PBS). Samples were all kept under nitrogen gas beforeuse. Samples were approximately the same diameter as the conical tubeused to contain the experiment.

FIG. 6 illustrates stress vs. strain curve of PCL-CPO polymer withdifferent loading concentrations of CPO (A) pre-activation and (B)post-activation in PBS for 24 h. (C) average Young's modulus ofelasticity for samples pre-activation and post-activation as a functionof CPO loading concentrations. (D) maximum elongation for pre-activationand post-activation in PBS for 24 h.

FIG. 7 illustrates Force versus displacement for different honeycombstructures with inlaid pictures of each sample in each respective trial.Pore size is the dimensions of the hexagon (i.e. 5 mm pore sizecorresponds to mesh comprised of hexagonal holes with heights and widthsof 5 mm). The separation between pores for all three sizes stayedconsistent at 3 mm.

FIG. 8 illustrates Scanning electron microscopy (SEM) images of thinfilm samples at different concentrations and multiple magnificationsboth before and after immersion in PBS for 24 hours. Thin films with twoconcentrations of CPO were imaged, specifically the 5 wt % film at (A)100× magnification and (B) 1000× magnification, the 15 wt % film beforeimmersion at (C) 100× magnification and (D) 1000× magnification. The 5wt % film was imaged after immersion at (E) 100× magnification and at(F) 1000× magnification. The 15 wt % film was imaged after immersion at(G) 100× magnification and at (H) 1000× magnification.

FIG. 9 illustrates Cell cultures fixed with 4% formaldehyde andpermeabilized with 0.1% Triton X-100 after exposure to differentconcentrations of CPO mesh at both time periods. (A) 0 wt % at 2 days;(B) 0 wt % at 6 days; (C) 5 wt % at 2 days; (D) 5 wt % at 6 days; (E) 15wt % at 2 days; (F) 15 wt % at 6 days. (G) shows the percent viablecells for the different concentrations at both time points. N=300 inthree biological replicates. Magnification 40×; Objective N/A: 0.65.

FIG. 10 illustrates Immunofluorescence images of HMS 32 cells showphalloidin staining (green) to show F-actin distribution in thecytoplasm and DAPI to stain nuclei (blue) on (A) control (no film) and(B) 15 wt % CPO in vial on day 6 of cell culture. (C) Bar graphs withpercentage of viable cells stained positive for caspase-3 on day 6 ofculture in control and 15 wt % CPO vials. n=150 in three technicalreplicates. Magnification 40×; Objective N/A: 0.65. *p<0.05; ***p<0.01;***p<0.001.

DETAILED DESCRIPTION

For the purposes of promoting an understanding of the principles of thepresent disclosure, reference will now be made to the embodimentsillustrated in the drawings, and specific language will be used todescribe the same. It will nevertheless be understood that no limitationof the scope of this disclosure is thereby intended.

In the present disclosure the term “about” can allow for a degree ofvariability in a value or range, for example, within 10%, within 5%, orwithin 1% of a stated value or of a stated limit of a range.

In the present disclosure the term “substantially” can allow for adegree of variability in a value or range, for example, within 90%,within 95%, or within 99% of a stated value or of a stated limit of arange.

The present disclosure provides a cost effective, scalable, andsolvent-free process of manufacturing polycaprolactone (PCL)-calciumperoxide (CPO) resorbable thin film as an alternative to surgical meshwith ability to provide effective mechanical strength and sustainedrelease of oxygen for limiting hypoxia-induced necrosis within theimplanted area (FIG. 1). The utilized PCL polymer matrix is advantageousbecause of its low melting temperature (−55° C., depending on themolecular weight), extended time required to fully degrade,biocompatibility, and lower cost as compared to biological polymers.Previous studies have shown PCL to last for more than two years as animplant in the body which is ideal for the required duration of servicein many implanted surgical meshes. Furthermore, the current disclosurehas utilized manufacturing methods take advantage of series of scalablemanufacturing methods including hot melt mixing, compression molding,and laser micro machining. The demonstrated fabrication procedure can bereadily adapted to roll-to-roll fabrication processes forcommercialization and mass production.

Materials and Methods

PCL-CPO Melt Mixing

It is necessary to have a uniform mixture before compression molding inorder to have a consistent concentration of CPO throughout the filmafter molding. Therefore, PCL and CPO were melt mixed at differentweight ratios (0, 1, 5, and 15 wt %) in a beaker uniformly heated by awater bath on a hot plate set at 90° C. For faster melting, the PCLpellets (M_(w) 65,000, Sigma Aldrich) were gradually added in 1 gram ata time, until the final amount of PCL was melted. Next, under continuousstirring, proportional quantities CPO was added to the melted PCL toobtain the desired loading concentrations. The mixtures were stirredcontinuously for about 30 min to achieve uniformity. Finally, thePCL-CPO melt mixture was transferred onto a silicone sheet to solidifyand cool down to room temperature. All solid polymer mixtures werestored under nitrogen gas before the compression molding process (FIG.2, Step A).

Compression Molding

Approximately 2 grams of the PCL-CPO mixture was placed in between twoKapton sheets. The polymer mixture was then flattened between 2 metalsheets using a compression molding machine (Carver Hydraulic Unit Model#3925). Prior to applying pressure, the mixture was heated at 82.2° C.for 20 minutes. 11 tons of pressure was then applied for 5 minutes. TheKapton sheets were removed from the compression molder and allowed tocool for another five minutes before carefully removing the thin filmfrom the Kapton sheets. All thin films were stored in a nitrogen boxbefore further laser processing and other characterization (FIG. 2, StepB and Step C).

Laser Micromachining

Desired surgical meshes with different dimensions of hexagonal shapedpores were generated using CorelDraw (Corel Corporation) cut in to thePCL-CPO thin films using a CO₂ laser cutting and engraving system(PLS6MW, Universal Lasers, Inc., Scottsdale, Ariz.). For clean cuts thelaser source power and scanning speed was set to 2 m/s and 60 W,respectively. FIG. 2, Step D and Step E).

Meshes with three different hexagonal pore dimensions were laser cutincluding: 2.5 mm height×2.5 mm width, 5 mm height×5 mm width, and 10 mmheight×10 mm width. The hexagons for each respective pore size wereevenly spaced 3 mm apart from each other. FIG. 3 illustrates thecomplete fabrication process for the functional PCL and CPO mesh.

Oxygen Generation Measurements

Oxygen generation capabilities of the composite meshes with differentratios of CPO filling content were determined by submerging laser cut 1cm in diameter circular thin film mesh samples into a 50 mL conical tube(Corning, USA) containing 1.5 mL of phosphate buffered saline. Duringentire measurement the tubes were sealed from the atmosphericenvironment to minimize the escape of generated oxygen from thecontainer. The levels of dissolved oxygen in the container were measuredby a RedEye Oxygen Sensor patch (Lightwind, USA) placed inside thecontainer. The patch contains immobilized ruthenium complex dyes thatchange in fluorescence properties in the presence of oxygen. The changesin fluorescence properties of the patch were measured by a bifurcatedoptical probe (RE-BIFBORO-2) pointed at the part of the container wherethe RedEye patch was located. This eliminates the need to open thecontainer for the oxygen measurements.

Tensile Testing

It is important to understand the mechanical properties of the developedmesh because there are strength requirements for mesh to provideadequate structural support in the body. In order to understand themechanical properties of the mesh, we performed tensile testing upon themesh according to the following parameters. Rectangular samples (3 mm×35mm) were fabricated using a laser cutter equipped with a CO₂ laser withsame setting as described before. The samples were tested using an ADMETMtestquattro universal tensile machine at room temperature with a jograte of 3 mm/min.

It is also important to understand the mechanical properties of the meshafter exposure to the conditions of the body. Therefore, tensile testingwas also performed upon samples that had been submerged in phosphatebuffered saline (PBS) for 24 hours under the same parameters asdescribed previously.

Mesh Tensile Testing

Tensile testing was also performed upon samples that were processed intomesh patterns in order to both determine mechanical properties of themesh and acquire data to build a model for the material. The tensiletesting was done on a stretching machine capable of applying forceunilaterally upon both ends of a sample. A jog rate of 3 mm/min was usedfor the tests.

Differential Scanning Calorimetry (DSC) and Thermogravimetric Analysis(TGA)

In addition to mechanical properties, it is also useful to understandthe thermal properties of a material. It is important to ensure thematerials being implanted in the body will neither degrade nor deformdue to the temperatures of the environment. DSC and TGA measure theglass transition temperature, melting temperature and degradationtemperature.

About 5 mg of each sample was used for TGA. The samples were analyzedusing a TG 209 F3 Tarsus. CPO and the 15 wt % thin film were heated to amax temperature of 900° C., while pure PCL was heated to a maxtemperature of 800° C. All samples were heated at a rate of 20° C./minunder a flow of nitrogen gas.

For DSC, 2-10 mg of each sample was analyzed using a DSC 214 Polyma. Thesamples were all heated to a temperature of 200° C. at a rate of 20°C./min, cooled to 25° C., then heated again to 200° C., all under a flowof nitrogen gas. The purpose of the first heating of the samples was toremove any impurities from the samples.

Scanning Electron Microscopy (SEM)

In order to visually confirm the presence of CPO and inspect themorphology, SEM was performed upon the thin film samples. The sampleswere all sputter coated (SPI-Module, SPI Supplies, West Chester, Pa.)with Au—Pd for 3 minutes at 1-minute intervals. Scanning electronmicroscopy (field-emission SEM, Hitachi S-4800) operated at 20.0 kV at100× and 1000× magnification was used to assess the morphology andmicrostructures of the polymer composite films before and after exposureto liquid environment.

Cell Culture Biocompatibility with Human Mammary Cells

It is essential to ensure biocompatibility with the thin film because itwill be implanted inside the body. Therefore, human mammary stromalfibroblasts, HMS-32 hTERT cells (between passages 4 and 8) were used toperform cell culture assays with thin films of different concentrations.HMS 32 cells were cultured in Dulbecco modified Eagle medium (DMEM)/F12medium (Invitrogen Inc., Carlsbad, Calif.) supplemented with additives,insulin (250 ng/ml; Sigma-Aldrich), hydrocortisone (500 ug/ml; BDBiosciences, San Jose, Calif.), sodium selenite (2.6 ng/mL; BDBiosciences), transferrin (10 ug/mL; Sigma-Aldrich), transforming growthfactor beta 1(TGFβ1) (30 pg/ml; Thermofisher, Waltham Mass.), andfibroblast growth factor (FGF)) 5 ng/ml (Thermofisher). The cells wereplated on 18 mm coverslips (Thermofisher, Waltham Mass.) with seedingdensity of 2500 cells/cm² and placed in a 12-well plate with and withoutcircular PCL-CPO strips (10 mm diameter) of the followingconcentrations: 0 (control), 5, and 15 wt % CPO. Cells were cultured at37° C. in a humidified environment with 5% carbon dioxide, and theculture medium was changed every two days. Cells on days 2 and 6 ofculture, using caspase-3 staining cell survival and apoptosis, weremonitored under different culture conditions.

Hypoxic Cell Culture Viability

In addition to biocompatibility of the thin films, it is important tomake sure the thin films support the viability of cells in hypoxicconditions. To accomplish this, HMS-32 cells were cultured on 12 mmglass cover slips (Thermofisher) as described above and placed in 20 mLscintillation vials without (control) and with a strip of 15 wt % CPOthin film. Hypoxic conditions were introduced into the vials by bubblingnitrogen gas into the solution for 15 minutes and sealing the vialsthereafter until day 6. The seal was not removed and there was no mediumchange. Cells were exposed to low CO₂ and O₂ levels. On day 6, cells on12 mm cover slips were processed for staining with Caspase-3 to assesscell viability.

Fluorescence Immunostaining

Antibody against Caspase-3 (Cell signaling technologies, Boston, Mass.,1:400 dilution) was used for testing the toxicity of the PCL-CPOmaterial at different concentrations and to assess survival in thepresence of the film under hypoxic conditions. Cells were fixed with 4%paraformaldehyde (Sigma-Aldrich) and processed for immunostaining asdescribed previously. Cells were stained with Alexaflour @480 conjugatedphalloidin (Thermofisher; 1:40 dilution) to stain F-actin (green) in thecytoplasm. Nuclei were counter-stained with DAPI (500 μg/ml). Imageswere recorded using Q-capture image acquisition software linked to aBXI70 inverted fluorescence microscope (Olympus, Waltham, Mass.), with40× objective (NA=0.65).

Results and Discussion

In order to assess the process viability of the thin film, DSC and TGAwere performed first to ensure the temperature required to melt PCL in awater bath would not surpass the degradation temperature of the polymerand the temperature in the body would not cause deformation ordegradation of the mesh. As shown in FIG. 4A, the peak decompositiontemperatures (339° C., 397° C., and 296° C.) for PCL, CPO and 15 wt %thin film, respectively, are more than 240° C. above the peak meltingtemperatures of the 15 wt % thin film and pure PCL (51.9° C. and 55.9°C., respectively), shown in FIG. 4B. The temperatures at which the thinfilm either melt or degrade are far higher than temperatures inside thebody. Therefore, there is no risk for the mesh to either deform ordegrade due to excessive temperatures. Additionally, becausemanufacturing temperatures do not exceed, 90° C., there is no risk fordegradation during fabrication.

After confirmation of thermal stability, oxygen generation capacitymeasurements were performed to quantify the amount of oxygen produced byeach concentration of CPO. FIG. 5 shows the results from the oxygengeneration experiments. All thin films experienced a short burst ofoxygen production immediately when the films were placed in the PBS,presumably from exposed CPO on the surface of the thin film, before adelay of no oxygen production ranging from 18 minutes for 1 wt % to 47minutes for 15 wt %. As the concentration of CPO increased, the amountof oxygen produced also increased (FIG. 5A). The maximum levels ofoxygen produced from the thin films ranged from the 1 wt % CPO thin filmwith 22.1% oxygen to the 15 wt % CPO thin with 25.5% oxygen (FIG. 5B).After the maximum oxygen levels were reached, there is a decline inconcentration of oxygen for all samples containing CPO. We believe thedecrease in oxygen levels after reaching the peak levels is a result ofthe oxygen levels equilibrating in the container. Because the thin filmis placed very close to the RedEye sensor, most of the oxygen isgenerated near the sensor, hence a higher concentration near the sensor.Therefore, the local concentration of oxygen near the sticker must behigher than the surrounding liquid until the oxygen diffuses equallythroughout the apparatus.

In thin films with lower amounts of CPO (1 and 5 wt % CPO), a plateau isobserved because there is less CPO to react in these thin films, and forthat reason they produce less oxygen than the 15 wt % thin film. Thelower oxygen output from the 1 and 5 wt % thin films would thereforeequilibrate sooner than the 15 wt % thin film.

FIG. 6A shows the tensile properties of the thin films with varyingamounts of CPO prior to being placed in a wet environment. The resultsfrom the experiment show the tensile strength of the thin filmsincreasing as the concentration of CPO in the thin films increases, from12900 kPa (±2170) for the 1 wt % thin film to 14500 kPa (±695) for the 5wt % thin film to 15900 kPa (±546) for the 15 wt % thin film. Theexception to this trend is the pure PCL thin film which has a tensilestrength of 15500 kPa (±1770).

FIG. 6B shows the tensile properties of the thin films after beingsoaked in PBS for 24 hours. The thin films with CPO decreased in tensilestrength as the concentration of CPO increased, from 14900 kPa (±449) inthe 1 wt % thin film to 12300 kPa (±1490) in the 5 wt % thin film to11800 kPa (±1330) in the 15 wt % thin film. However, the pure PCL thinfilm exhibited a much higher tensile strength than any of the otherfilms at 24200 kPa (±2410). The surface porosity created as the CPOinitially present on the surface reacts most likely creates weaker areasin the thin films that lessens the tensile strength of the thin filmswith CPO. However, PCL thin film without CPO do not experience thedecrease in tensile strength.

FIG. 6C illustrates the differences in the amount of elongation the thinfilms can endure pre-activation and post-activation. The maximumelongation decreases as the amount of CPO in the thin films increase forboth pre-activation and post-activation. This decrease in elongation isalso due to the surface porosity created by CPO reacting with PBS, whichcauses a more brittle material.

The Young's modulus of Elasticity, shown in FIG. 6D, shows a largedifference between samples before and after activation (before and afterimmersion, respectively). For the 0 and 1 wt % CPO samples beforeimmersion in PBS, the Young's modulus of elasticity is less than thesamples previously immersed in PBS, whereas the opposite is true for the5 and 15 wt % CPO samples. It appears from FIG. 6D that there is acritical amount of CPO in PCL where the Young's modulus of elasticity isnot affected by soaking in PBS for 24 hours because of this observedtrend. It is known that water breaks down the polymer chains in PCL,leading to a more brittle material, which explains the behavior of thepure PCL samples. For the samples with larger amounts of CPO, the higherdensity of pores created on the surface of the material from CPOreacting with water must compromise the structural integrity of thematerial more than the water makes the polymer matrix brittle, resultingin a decrease in Young's modulus for samples post-activation. Thiscritical point must be somewhere in between 1-5 wt % CPO.

FIG. 7 shows the force required to break laser machined meshes withdifferent pore sizes onto thin films PCL with 15 wt % CPO filler. Theseresults are supported by the simulations done in Abaqus. For each poresize (2.5 mm, 5 mm and 10 mm), failure was observed at the point wheremaximum von Mises stress occurred. Experimentally, the sample undergoesthe same amount of elongation as in simulation. As can be seen from theimages taken from simulation, maximum stress occurred mostly at theedges of the mesh due to the stress concentration at the corners of thesample. These results are supported by the simulations done in Abaqusshown in FIG. 7. The experimental results are all very close or equal tothe simulation results, thus indicating an accurate model.

SEM images were conducted to visually confirm the uniformity of the CPOparticles distribution in the thin film composite. FIG. 8A shows the 5wt % CPO thin film at 100× magnification. FIG. 8B shows the same film at1000× magnification. The CPO particles can be identified as spots ofwhite and lighter shades of grey. Although the CPO in the 5 wt % thinfilm is distributed evenly, there is not a high concentration (FIG. 8B).FIG. 8C and FIG. 8D show the 15 wt % thin film at 100× and 1000×magnification, respectively. It can be seen clearly in FIG. 8D that theCPO particles are still distributed evenly but in a higher concentrationthan in the 5 wt % thin film. Additionally, it is evident in FIG. 8B andFIG. 8D that some of the CPO particles are embedded beneath the surfacebecause some of the CPO spots are light grey rather than white,signifying their presence just below the surface of the PCL. Thispartial encapsulation of CPO particles allows a more sustained releaseof oxygen as the PCL breaks down in the body rather than a burst releaseif all the CPO was on the surface.

SEM images were also taken of the composite films after being immersedin PBS to assess its change in structural properties. FIG. 8E and FIG.8F show the 5 wt % thin film after being soaked in PBS for 24 hours.FIG. 8F shows the formation of a hole in the surface of the polymer filmwhere a particle of CPO was before reacting with PBS. FIG. 8G and FIG.8H show the 15 wt % film after being soaked in PBS for 24 hours. Highermagnification images confirm that the increase in CPO filler contentleaved behind more holes in the surface of the polymer film afterimmersion into PBS solution, FIG. 8H.

To assess the biocompatibility of the thin films with human cells,different concentrations of thin films were cultured with HMS-32 cells.Based on caspase-3 based cell viability measurement on the HMS-32 cells,the PCL-CPO is biocompatible even at 15 wt %, the highest concentrationtested. The cell survival did not change significantly as compared tothe control (FIG. 9G). In this case we have used an invertedfluorescence microscope to determine the cell viability and subsequentimaging. Based on caspase staining, 90-92% cells were viable whencultured on any of the substrate materials (FIG. 9G). There was nosignificant difference in percentage of viable cells on day 2 which isan earlier stage of assessing the compatibility and on day 6 by whichfibroblasts are expected to differentiate in the presence of growthadditives provided (FIG. 9A). There was no change in the phenotype ofthe cells (FIG. 9A-FIG. 9F). These results show that the presence ofincreasing amounts of CPO do not greatly impact the viability of cells.Because of the viability results, we chose to focus on the 15 wt % thinfilm, as it produces the most oxygen of all the thin films.

Although the films were determined to be biocompatible with human cells,it was still necessary to show the capacity of the films to aid in thesurvival of cells in hypoxic conditions. For this test, PCL film with 15wt % CPO fillers were introduced into sealed cell culture vial inhypoxic conditions. To assess the oxygen-generating the functionality oncell survival, the amount of live or dead cells was visualized andquantified after day 6, FIG. 9A-FIG. 9B. The control cell culture groupin hypoxia condition underwent significant amounts of cell death withonly 15% surviving cells. However, the PCL film with 15 wt % CPOdemonstrated cell survival rates of up to 80%. Although this number issignificantly lower than usual survival rate in HMS-32 cells culturedunder regular conditions, it is expected that there is more cell deathin the vials with low 02 and CO₂. Because of the increased survivalrates for cells in hypoxic conditions exposed to the 15 wt % thin film,we believe PCL-CPO thin film composite meshes show great promise for usein surgical mesh applications.

Oxygen-generating biodegradable surgical mesh were fabricated byincorporating varying concentrations of CPO in PCL polymer matrix. Theuse of scalable manufacturing methods, including hot melt mixing,compression molding, and laser processing enabled to greatly reduce themanufacturing complexly and production of the composite mesh. Thetensile strength of the composites with CPO filler of up to 15 wt % wasfound to be provide enough strength for a wide range of hernioplastiesapplication with the require a tensile strength of 24 kPa. The rate ofoxygen-generation of the polymer composite upon exposure to liquidconfirmed a consistent increase with higher CPO filler content.Cytotoxicity studies with human mammary stromal fibroblasts confirmedthe biocompatibility of a polymer composite with different CPOconcentrations of up to 15 wt %. In vitro cell viability studies inhypoxic conditions confirmed the ability of the PCL composite with 15 wt% CPO filler to effectively reduce hypoxia-induced cell death bylimiting the necrosis over 6 days of culture. The demonstrated PCL-CPOcomposite shows great potential to replace or supplement currentlyavailable surgical meshes. Future work should investigate long-termimpacts that the thin film meshes may have on the body.

In one embodiment, the present disclosure provides a surgical mesh,wherein the surgical mesh has a flexible basic structure and comprises aplurality of pores, wherein the surgical mesh has a first face and asecond opposite face, wherein the surgical mesh is made of asubstantially homogeneous material comprising a biodegradable polymericmaterial and an oxygen-generating material.

In one embodiment regarding the surgical mesh of this disclosure,wherein the biodegradable polymeric material may be but is not limitedto poly(lactic-co-glycolic acid) (PLGA), thermoplastic polyurethane(TPU), or polycaprolactone (PCL). In one aspect, the biodegradablepolymeric material is polycaprolactone (PCL).

In one embodiment regarding the surgical mesh of this disclosure,wherein oxygen-generating material may be but is not limited to sodiumpercarbonate, calcium peroxide, magnesium peroxide, or any combinationthereof. In one aspect, the oxygen-generating material is calciumperoxide.

In one embodiment regarding the surgical mesh of this disclosure,wherein the biodegradable polymeric material comprises polycaprolactone,wherein the oxygen-generating material comprises calcium peroxide.

In one embodiment regarding the surgical mesh of this disclosure,wherein the polycaprolactone has a weight percentage of 85% to 98% ofthe total weight of the surgical mesh, wherein the calcium peroxide hasa weight percentage of 2% to 15% of the total weight of the surgicalmesh.

In one embodiment regarding the surgical mesh of this disclosure,wherein the plurality of pores have uniform size and shape to provideconsistent oxygen releasing rate.

In one embodiment regarding the surgical mesh of this disclosure,wherein the surgical mesh has a thickness of 100-500 μm, 100-400 μm,100-300 μm, 100-200 μm, 100-150 μm.

In one embodiment regarding the surgical mesh of this disclosure,wherein the surgical mesh is biocompatible and has over 80% of cellsurvival rate in hypoxic condition.

In one embodiment, the present disclosure provides a method of preparingthe surgical mesh of the present disclosure, wherein the methodcomprises:

providing a biodegradable polymeric material and melting thebiodegradable polymeric material;

providing an oxygen generating material and adding the oxygen generatingmaterial to the melted biodegradable polymeric to form a substantiallyhomogenous melt mixture;

cooling the substantially homogenous melt mixture to provide asubstantially, homogenous solid mixture;

providing an amount of the substantially homogenous solid mixture andheating to an elevated temperature and then compression molding thesubstantially homogenous solid mixture to a sheet with a thickness of100-500 μm; and laser micromachining the sheet to generate a pluralityof pores.

In one embodiment regarding the method of preparing the surgical mesh ofthe present disclosure, wherein the biodegradable polymeric materialcomprises polycaprolactone, wherein the oxygen-generating materialcomprises calcium peroxide.

In one embodiment, the present disclosure provides a method of using thesurgical mesh of the present disclosure, wherein the method comprises:

providing the surgical mesh of claim 1 to a patient during a surgeryprocess; and

applying the surgical mesh to a surgical site of the patient to providestructural support to the surgical site, and to provide an oxygensource.

In one embodiment regarding the method of using the surgical mesh of thepresent disclosure, Wherein the method is used to control bacteriaprefiltration.

In one embodiment regarding the method of using the surgical mesh of thepresent disclosure, wherein the biodegradable polymeric materialcomprises polycaprolactone, wherein the oxygen-generating materialcomprises calcium peroxide.

Those skilled in the art will recognize that numerous modifications canbe made to the specific implementations described above. Theimplementations should not be limited to the particular limitationsdescribed. Other implementations may be possible.

We claim:
 1. A surgical mesh, wherein the surgical mesh has a flexiblebasic structure and comprises a plurality of pores, wherein the surgicalmesh has a first face and a second opposite face, wherein the surgicalmesh is made of a substantially homogeneous material comprising abiodegradable polymeric material and an oxygen-generating material. 2.The surgical mesh of claim 1, wherein the biodegradable polymericmaterial comprises poly(lactic-co-glycolic acid) (PLGA), thermoplasticpolyurethane (TPU), or polycaprolactone (PCL).
 3. The surgical mesh ofclaim 1, wherein the oxygen-generating material comprises calciumperoxide.
 4. The surgical mesh of claim 1, wherein the biodegradablepolymeric material comprises polycaprolactone, wherein theoxygen-generating material comprises calcium peroxide.
 5. The surgicalmesh of claim 4, wherein the polycaprolactone has a weight percentage of85% to 98% of the total weight of the surgical mesh, wherein the calciumperoxide has a weight percentage of 2% to 15% of the total weight of thesurgical mesh.
 6. The surgical mesh of claim 1, wherein the plurality ofpores have uniform size and shape to provide consistent oxygen releasingrate.
 7. The surgical mesh of claim 1, wherein the surgical mesh has athickness of 100-500 μm.
 8. The surgical mesh of claim 1, wherein thesurgical mesh is biocompatible and has over 80% of cell survival rate inhypoxic condition.
 9. A method of preparing the surgical mesh of claim1, wherein the method comprises: providing a biodegradable polymericmaterial and melting the biodegradable polymeric material; providing anoxygen generating material and adding the oxygen generating material tothe melted biodegradable polymeric to form a substantially homogenousmelt mixture; cooling the substantially homogenous melt mixture toprovide a substantially homogenous solid mixture; providing an amount ofthe substantially homogenous solid mixture and heating to an elevatedtemperature and then compression molding the substantially homogenoussolid mixture to a sheet with a thickness of 100-500 μm; and lasermicromachining the sheet to generate a plurality of pores.
 10. Themethod of claim 9, wherein the biodegradable polymeric materialcomprises polycaprolactone, wherein the oxygen-generating materialcomprises calcium peroxide.
 11. A method of using the surgical mesh ofclaim 1, wherein the method comprises: providing the surgical mesh ofclaim 1 to a patient during a surgery process; applying the surgicalmesh to a surgical site of the patient to provide structural support tothe surgical site, to provide an oxygen source, and to control bacteriaprefiltration.
 12. The method of claim 11, wherein the biodegradablepolymeric material comprises polycaprolactone, wherein theoxygen-generating material comprises calcium peroxide.